Optical Coherence Tomography (OCT) is an emerging non-invasive biomedical imaging technology that can perform cross-sectional imaging of tissue microstructures in vivo and in real-time. OCT is analogous to ultrasound, except that it uses low coherence light, rather than acoustic waves. The echo delay time or the depth of light backscattered from tissue is measured using a technique referred to as low coherence interferometry.
OCT has significant advantages over other medical imaging technologies. Medical ultrasound, magnetic resonance imaging (MRI), and confocal microscopy are ill suited to high-resolution morphological deep tissue imaging, as ultrasound and MRI have insufficient resolution for imaging microstructures, while confocal microscopy lacks the ability to image deeply enough (i.e., beyond several hundred micrometers in highly scattering tissues), an ability that is required for morphological tissue imaging. OCT is analogous to ultrasound B-mode imaging, except that it uses low coherency, near-infrared light, rather than sound, and no matching medium is required. OCT imaging is non-invasive, and imaging can be performed in situ and in real time. In addition to micro-structural imaging, OCT can also provide additional functional information, such as high-resolution Doppler flow, and spatially revolved tissue spectroscopy.
As indicated above, a fundamental aspect of OCT is the use of low coherence interferometry (either in the time domain or the Fourier domain). In conventional laser interferometry, the interference of light occurs over a distance of meters. In OCT, the use of broadband light sources (i.e., light sources that can emit light over a broad range of frequencies) enables the interference to be generated within a distance of micrometers. Such broadband light sources include super luminescent diodes (i.e., super bright light emitting diodes (LEDs)), extremely short pulsed lasers (i.e., femto-second lasers) and wavelength/frequency-swept lasers. White light can also be used as a broadband source.
Essentially, the combination of backscattered light from the sample arm and reference light from the reference arm gives rise to an interference pattern, but only if light from both arms have traveled “substantially the same” optical distance (where “substantially the same” indicates a difference of less than a coherence length). By scanning the mirror in the reference arm or using Fourier domain techniques, a reflectivity profile of the sample can be obtained. Areas of the sample that reflect more light will create greater interference than areas that reflect less light. Any light that is outside the short coherence length will not contributes significantly to the interference signal. This reflectivity profile, referred to as an A-scan, contains information about the spatial dimensions and location of structures within the sample. An OCT image (i.e., a cross-sectional tomograph generally referred to as a B-scan), may be achieved by laterally combining multiple adjacent axial scans at different transverse positions.
FIG. 1 (Prior Art) schematically illustrates a conventional OCT system. This system includes a Michelson interferometer that uses a low coherence light source 20. The light source is coupled to an OCT probe 24 in the sample arm and to a reference arm 28 through an optical fiber coupler or beam splitter 22. The sample arm delivers an optical beam from the light source to a target 26 (generally a tissue sample) and collects the backscattered light. The reference arm provides a reference optical path length for the interference signal (with the path length scanned as in time-domain OCT or unscanned as in Fourier domain OCT). Path length scanning can be achieved, for example, by using a translating retro-reflective mirror or a phase-controlled scanning delay line (not separately shown). A backscattered intensity versus depth data set is developed with an axial scan. Two- or three-dimensional data sets formed by multiple adjacent axial scans are obtained by scanning the OCT beam along the transverse direction after each axial scan. A photodetector 30 (or a detector array) produces a corresponding analog signal comprising the data set. The analog signal is processed by detection electronics module 32, which produces corresponding digital data. The resulting digital data set can be further processed, displayed and stored, using a computer 38, as a false-color or gray-scale map, to form a cross-sectional OCT image. Barrett's esophagus is a chronic metaplastic condition characterized by a change in the epithelial lining of the esophagus, from normal squamous epithelium to a specialized columnar epithelium. Its prevalence is highly correlated to gastro esophageal reflux disease. Although Barrett's esophagus often does not cause symptoms, individuals having this condition have a significantly higher risk of developing esophageal cancer. The incidence of this usually lethal malignancy has increased 350% over the past two decades in the United States. Currently, the standard surveillance procedure for Barrett's epithelium is endoscopy, along with four-quadrant pinch biopsies at 1-2 cm intervals throughout the Barrett's epithelium. Recently, minimally invasive endoscopic ablative therapies (such as PDT, and electro or laser coagulation) have been developed and show significant potential to eradicate Barrett's esophagus, with the goal of preventing the development of esophageal cancer. Following ablation, there is return of normal-looking squamous epithelium. However, biopsy studies have shown that residual Barrett's esophagus remains underneath the neo-squamous epithelium in approximately 5% of cases. The true prevalence of this condition could be much higher, considering the random sampling nature of biopsy. Residual Barrett's epithelium under squamous epithelium has been reported to lead to the development of cancer. A major concern for ablative therapies is that if Barrett's remains under squamous epithelium, it cannot be detected using conventional endoscopic technologies. Currently, there are no imaging techniques that are capable of detecting islands of Barrett's epithelium under squamous epithelium. It would thus be desirable to provide a method and apparatus that can be used to detect such residual Barrett's epithelium.
Referring to FIG. 1, it will immediately be apparent that the form factor or size of OCT probe 24 limits the application of this technology within the body of a patient. However, reduced form factor OCT probes have been developed for use with intra-luminal, intravascular and interstitial catheters. FIG. 2A schematically illustrates a Prior Art catheter-based OCT probe 40, which includes a single mode optical fiber 44 for conveying light to and from a sample 52, a gradient index (GRIN) lens 46, and a beam deflecting element 48. These optical elements are disposed in a housing 42, and an opening 50 in housing 42 enables light to reach sample 52. In general, an object distance, L1, combined with properties of the lens, determines a working distance, W1, that can be achieved.
FIG. 2B (Prior Art) graphically illustrates the relationship between object distance, working distance, and transverse resolution that can be achieved using the probe configuration of FIG. 2A. At a working distance of about 8 mm (required for esophageal imaging), the best transverse resolution obtainable with a given GRIN lens using the probe configuration of FIG. 2A is about 75 microns, assuming that that the optical fiber and GRIN lens are coupled using optical cement (i.e., L1 corresponds to the thickness of the layer of optical cement). Unfortunately, that resolution is insufficient to enable fine structures (i.e., structures less than about 50 um in size) to be detected using OCT imaging. It would be desirable to provide additional probe designs facilitating improved resolution, while maintaining a compact form factor.